MEMS affinity sensor for continuous monitoring of analytes

ABSTRACT

Techniques for monitoring a target analyte in a sample using a polymer capable of binding to the target analyte are disclosed. A microdevice useful for the disclosed techniques includes a semi-permeable membrane structure, a substrate, a first and second microchambers formed between the membrane structure and the substrate. The first microchamber can be adapted to receive a solution including the polymer, and the second microchamber can be adapted to receive a reference solution. Environmental target analyte can permeate the semi-permeable membrane structure and enters the first microchamber and the second microchamber. Based on the difference in a property associated with the polymer solution that is responsive to the target analyte-polymer binding, and the corresponding property associated with reference solution, the presence and/or concentration of the target analyte can be determined.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of International Patent ApplicationNo. PCT/US12/048819, filed Jul. 30, 2012, which claims priority fromU.S. Provisional Application No. 61/513,335, filed Jul. 29, 2011; U.S.Provisional Application No. 61/538,732, filed Sep. 23, 2011; U.S.Provisional Application No. 61/542,113, filed Sep. 30, 2011; and U.S.Provisional Application No. 61/542,139, filed Sep. 30, 2011. Thedisclosure of each of the foregoing applications is herein incorporatedby reference by its entirety.

STATEMENT REGARDING FEDERALLY FUNDED RESEARCH

This invention was made with government support under grant numberDK63068-05, awarded by the National Institutes of Health, and grantnumber ECCS-0702101, awarded by the National Science Foundation. Thegovernment has certain rights in this invention.

BACKGROUND

Diabetes mellitus is a metabolic disease characterized by persistenthyperglycemia (high blood sugar levels). Complications induced bydiabetes, such as heart disease, stroke, hypertension, blindness, kidneyfailure, and amputation deprive the lives of 231,404 people in Americaas recently as 2007, making diabetes the seventh leading cause of death.

Glucose monitoring can reduce the occurrence rate and severity ofcomplications caused by hyperglycemia or hypoglycemia. Thus, it isimportant to closely monitor abnormal blood sugar levels in diabetespatients so timely treatments (e.g., insulin injection, exercise, anddiabetic diet, intake of carbohydrate) can be administered. This can beachieved by continuous glucose monitoring, which involves eithernoninvasive or minimally invasive detection of glucose. Noninvasivemethods can extract interstitial fluid (ISF) glucose from the skin inminimally destructive approaches or measure blood glucose in contactlessmanners. Although noninvasive methods can be used for CGM,interferences, such as the complexity of skin structures, sweating,temperature, and exercise can impact the accuracy and reliability of thesystem, limiting their practical applications in CGM. Certain minimallyinvasive methods can use subcutaneous sensor implantation to monitor theglucose levels in ISF. In the steady state, ISF glucose concentration isidentical to that in blood. However, when the blood glucose levelsundergo rapid changes, time lags between the blood and the ISF glucoseconcentrations can occur.

Electrochemical glucose sensors, which use O₂ and H₂O₂ as the mediators,can also be subject to errors induced by fluctuations of oxygen levels.In addition, redox-active species, such as ascorbic and uric acids, cancompromise the selectivity and the accuracy of the glucose sensors.Other devices utilize artificial mediators (e.g., ferro/ferricyanide,hydroquinone, and ferrocene) as alternatives to oxygen for electrontransfer. However, competition of oxygen with the artificial mediatorsand potential leaching and toxicity of these artificial mediators canhinder the in-vivo applications of these devices.

MEMS devices offer miniature sizes and rapid time responses, and aresuited for implantable or noninvasive glucose sensors. MEMS technologycan be used in developing electrochemical CGM sensors. MEMS affinityglucose sensors can use Con A, boronic-acid based monomers and polymers,and GBP as the glucose receptors and measure the glucose-induced changesin the properties of these receptors. For example, viscosity changes dueto the binding of Con A or boronic acid-based polymers with glucose canbe exploited by optical or electrical detection of microcantilevervibration, piezoelectric detection of flow resistance, and hall effectdetection of microrotors.

There is a need to develop implantable glucose monitoring systems thatoffer improved long term accuracy and stability, low drift, resistanceto environmental parameter fluctuations, easier calibration, as well asthe capability of providing real-time report of a subject's glucoselevel via wireless telemetry.

SUMMARY

The disclosed subject matter provides techniques for monitoring a targetanalyte in a sample using a polymer capable of binding to the targetanalyte. In one aspect, an exemplary microdevice includes asemi-permeable membrane structure, a substrate, first and secondmicrochambers formed between the membrane structure and the substrate,and a suspended element positioned to be spaced apart from thesubstrate. The first microchamber can be adapted to receive a solutionincluding the polymer. The second microchamber can be adapted to receivea reference solution for screening effects not caused by the targetanalyte. The semi-permeable membrane structure can be permeable to thetarget analyte and impermeable to the polymer, such that when the sampleis placed in contact with the semi-permeable membrane structure, thetarget analyte, if present in the sample, permeates the semi-permeablemembrane structure and enters the first microchamber and the secondmicrochamber, respectively. The polymer can be prevented from escapingfrom the first microchamber through the semi-permeable membranestructure. The microdevice can further include the polymer solution inthe first microchamber and the reference solution in the secondmicrochamber.

In some embodiments of the microdevice, the binding of the polymer withthe target analyte causes a change in the permittivity of the polymersolution. In alternative embodiments, the suspended element of each ofthe first microchamber and the second microchamber includes avibrational element, and the binding of the polymer with the targetanalyte causes a change in the viscosity of the polymer solution, whichin turn influences the vibration of the vibrational element in the firstmicrochamber. The vibration of the vibrational element can be actuableby an external magnetic field. For example, the vibrational element caninclude permalloy.

In some embodiments, each of the first microchamber and the secondmicrochamber further include a top electrode and a bottom electrode,respectively. The top electrode can be included in the suspendedelement, and the bottom electrode can be included in the substrate. Inparticular embodiments, the top electrode can be supported by at leastone post formed from the substrate. In particular embodiments, the topelectrode for each of the first microchamber and the second microchambercan also be perforated.

The disclosed subject matter also provides a microdevice for monitoringa target analyte in a sample using a polymer capable of binding to thetarget analyte. An exemplary microdevice includes a semi-permeablemembrane structure; a substrate; a microchamber formed between thesemi-permeable membrane structure and the substrate. The microchambercan be adapted to receive a solution including the polymer, and includea suspended element positioned to be spaced apart from the substrate.The suspended element can include a perforated electrode, which can besupported on one or more posts formed from the substrate. Thesemi-permeable membrane structure can be permeable to the target analyteand impermeable to the polymer, such that when the sample is placed incontact with the semi-permeable membrane structure, the target analyte,if present in the sample, permeates the semi-permeable membrane andenters the microchamber. The polymer can be prevented from escaping fromthe microchamber through the semi-permeable membrane structure. Themicrodevice can further include the polymer solution in themicrochamber.

In various embodiments of the disclosed microdevices, the polymerreversibly binds with the target analyte. The microdevices can beadapted to be implantable in a subcutaneous tissue of a subject, e.g., amammal or a human subject. The microdevices can include integratedmicroheaters and/or temperature sensors, and can be coupled with awireless interface for transmitting signals representing measurement ofthe target analyte.

The disclosed subject matter also provides methods of using suchmicrodevices. In one example, a method includes loading a solutionincluding a target-analyte sensitive polymer into a first microchamber,and a reference solution into a second microchamber; placing a sample incontact with a semi-permeable membrane structure such that a targetanalyte in the sample permeates the semi-permeable membrane structureand enters the first microchamber and the second microchamber,respectively; and determining a presence and/or concentration of thetarget analyte in the sample.

BRIEF DESCRIPTIONS OF THE DRAWINGS

FIGS. 1A and 1B are schematic side views of microdevices according tosome embodiments of the disclosed subject matter.

FIG. 2 is a schematic top view of a microdevice according to someembodiments of the disclosed subject matter.

FIG. 3A is a diagram showing the composition ofPoly(N-hydroxyethylacrylamide-ran-3-acrylamidophenylboronic acid)(PHEAA-ran-PAAPBA), a glucose-sensitive polymer.

FIG. 3B is a plot illustrating the specificity of the polymer shown inFIG. 3A toward glucose in the presence of other sugars.

FIGS. 4A-4F are diagrams illustrating a representative fabricationprocess for the microdevice depicted in FIG. 1A. (A) Bottom electrodedeposition. (B) Parylene deposition and sacrificial layer patterning.(C) Parylene deposition. (D) Moving electrode deposition and Permalloyelectroplating. (E) Additional parylene layer deposition and aluminummask deposition. (F) Sacrificial layer removal and diaphragm release.

FIGS. 5A-5B are micrographs of (A) a single vibrational diaphragm and(B) a MEMS differential glucose sensor.

FIG. 6 is an exemplary setup for in-vitro and in-vivo characterizationof a MEMS differential glucose sensor.

FIG. 7 is a plot illustrating the response of a MEMS differentialglucose sensor of the disclosed subject matter to different glucoseconcentrations.

FIG. 8 is a plot illustrating the time-dependent vibration amplitude ofa vibrational diaphragm (indicated by the sensor capacitance) of a MEMSdifferential glucose sensor of the disclosed subject matter as glucoseconcentration changes from 60 to 90 mg/dL, then back to 60 mg/dL.

FIG. 9 is a plot depicting a simulation model to characterize theglucose diffusion in a MEMS differential glucose sensor.

FIG. 10 is a plot showing the simulation result of the time-dependentglucose concentration on the surface of the vibrational diaphragm ofglucose sensor as modeled in FIG. 9.

FIG. 11 is a plot showing simulated time-dependent glucose concentrationon the surface of the vibrational diaphragm in response to additionalglucose concentration increases.

FIG. 12 is a plot showing a comparison of the capacitance output of aMEMS glucose sensor in changing temperature in single-module anddifferential measurements.

FIG. 13 is a plot showing a comparison of the capacitance output of aMEMS glucose sensor when the sensor is exposed to 90 mg/dL glucosesolution for approximately 5 hours.

FIG. 14A is picture showing implanting a MEMS glucose sensor in alaboratory mouse

FIG. 14B is a plot showing the differential capacitance change of theMEMS glucose sensor implanted in FIG. 14A as compared to readings from acommercial glucometer.

FIG. 15 is a plot showing Clarke error grid to assess the clinicalaccuracy of the estimated glucose value obtained from calibratingdifferential capacitance with reference glucose value.

FIG. 16 is a schematic of a single-module MEMS dielectric glucose sensoraccording to some embodiments of the disclosed subject matter.

FIGS. 17A-17F are a diagram depicting a representative fabricationprocess of the single-module MEMS dielectric glucose sensor shown inFIG. 16: (a) Gold layer deposition and patterning to form a bottom goldelectrode, and passivation of the electrode by Parylene; (b) Sacrificialphotoresist layer deposition and patterning; (c) Parylene deposition andgold layer deposition and patterning to form a perforated electrode; (d)Parylene passivation layer deposition; (e) SU-8 deposition andpatterning to form a diaphragm and a microchamber; (f) SU-8 patterning,sacrificial layer removal, and semi-permeable membrane bonding.

FIG. 18 are a micrograph image of a MEMS dielectric glucose sensor asfabricated by the process illustrated in FIGS. 17A-17F before packaging.

FIG. 19 is a diagram showing a representative setup and acapacitance/voltage transformation circuit for capacitance measurementof a MEMS dielectric glucose sensor according to some embodiments of thedisclosed subject matter.

FIG. 20 is a plot showing frequency-dependent equivalent capacitance ofa MEMS dielectric glucose sensor in the absence of glucose.

FIG. 21 is a plot showing the capacitance differences of the MEMSdielectric glucose sensor at various glucose concentrations as comparedwith the sensor capacitance in the absence of glucose.

FIG. 22 is a plot showing time-dependent capacitance of the MEMSdielectric glucose sensor at 100 kHz as the sensor responded to glucoseconcentration changes from 60 to 120 mg/dL, which was then reversed to60 mg/dL.

FIG. 23 is a plot showing the capacitance drift of the MEMS dielectricglucose sensor at 100 kHz over an extended time duration as the glucoseconcentration was held constant at 60 mg/dL.

FIG. 24 is a plot showing the response of the MEMS dielectric glucosesensor to glucose free polymer solutions.

FIG. 25 is a plot showing time-dependent capacitance of the MEMSdielectric glucose sensor for a glucose concentration change (at 100kHz).

FIG. 26 is a plot showing the capacitance drift of the MEMS dielectricglucose sensor at 100 kHz over an extended time duration as the glucoseconcentration was held constant at 100 mg/dL.

FIGS. 27A-27F are a diagram illustrating a representative process forfabricating a MEMS differential dielectric glucose sensor depicted inFIG. 1B: (A) Gold layer deposition and patterning to form bottom goldelectrodes, and passivation of the electrodes by parylene; (B)Sacrificial photoresist layer deposition and patterning; (C) Parylenedeposition and gold layer deposition and patterning to form perforatedelectrodes; (D) Parylene passivation layer deposition; (E) SU-8deposition and patterning to form diaphragms and microchambers; (F) SU-8patterning, sacrificial layer removal, and semi-permeable membranebonding.

FIGS. 28A-28B are images of a differential MEMS glucose sensor accordingto some embodiments of the disclosed subject matter: (A) before, and (B)after packaging.

FIG. 29 is a plot showing the frequency responses of thePHEAA-ran-PAAPBA polymer solution and PAA polymer solution at 0 mg/dLglucose concentration when each was loaded in the sensing chamber of thedifferential MEMS glucose sensor of FIG. 28.

FIG. 30 is a plot showing the capacitance change of the differentialMEMS glucose sensor filled with PHEAA-ran-PAAPBA solution at varyingglucose concentration with respected to the capacitance at 0 mg/dLglucose concentration.

FIG. 31 is a plot showing the time-dependent capacitance of thedifferential MEMS glucose sensor at 32 kHz as the sensor responded toglucose concentration changes from 50 to 100 mg/dL, which was thenreversed to 50 mg/dL.

FIG. 32 is a plot showing the capacitance of the differential MEMSglucose sensor in response to a sequence of glucose concentrations.

FIG. 33 is a plot showing the capacitance of the differential MEMSglucose sensor as compared to a single module MEMS glucose sensor overan extended time duration as the glucose concentration was held constantat 50 mg/dL.

FIG. 34 is a plot showing a comparison of sensor capacitance output inchanging temperature in single module and differential measurements.

FIG. 35A is a picture showing sensor implantation in a laboratory mouseFIG. 35B is a plot of differential capacitance change of the implantedsensor according to FIG. 35A during the initialization.

FIGS. 36A-36C are plots for differential sensor capacitance changes ascompared to readings from a commercial glucometer of test subject mouse(A) one, (B) two, and (C) three.

FIG. 37 is a plot of Clarke error grid to assess the clinical accuracyof estimated glucose values obtained from calibrating differentialcapacitance with reference glucose values.

FIG. 38A is a schematic diagram showing a wireless communication betweenan implantable sensor and an external reader according to someembodiments of the disclosed subject matter.

FIG. 38B is an image showing an external reader and a wireless interfaceused in an example of the disclosed subject matter.

FIG. 39 is a plot showing the results of an example for testing thebasic function of the wireless interface by changing the dielectricssandwiched between two electrodes from air to water.

DETAILED DESCRIPTION

The disclosed subject matter provides for devices and techniques tomonitor target analytes. More specifically, the disclosed subject matterprovides for MEMS-based sensors and systems that can be used forcontinuous analyte monitoring, including continuous glucose monitoring(CGM). As used in, the microdevices are also referred to as sensors.

As used herein, the term “analyte” is a broad term and is used in itsordinary sense and includes, without limitation, any chemical speciesthe presence or concentration of which is sought in material sample bythe sensors and systems disclosed herein. For example, the analyte(s)include, but not are limited to, glucose, ethanol, insulin, water,carbon dioxide, blood oxygen, cholesterol, bilirubin, ketones, fattyacids, lipoproteins, albumin, urea, creatinine, white blood cells, redblood cells, hemoglobin, oxygenated hemoglobin, carboxyhemoglobin,organic molecules, inorganic molecules, pharmaceuticals, cytochrome,various proteins and chromophores, microcalcifications, electrolytes,sodium, potassium, chloride, bicarbonate, and hormones. In oneembodiment, the analyte is glucose. In various embodiments, the analytescan be other metabolites, such as lactate, fatty acids, cysteines andhomocysteines.

As used herein, the term “suspended element” refers to a thin filmstructure suspended in the microchamber(s) of the microdevice. Thesuspended element can include subparts, e.g., a thin film electrode, apassivation layer. In some embodiments, the suspended element includes amagnetically active component that can move in response to an externalmagnetic field.

As used herein, the term “vibrational element” refers to a mechanicalmoving part, which is capable of vibrating. The vibrational element asused in presently disclosed subject matter includes, but is not limitedto, a vibrational diaphragm. The vibrational element can be, or a partof the suspended element.

The microdevices of the disclosed subject matter can be eitherviscosity-based or permittivity based. FIG. 1A illustrates the structureof an example viscosity-based microdevice. As shown in FIG. 1A, themicrodevice includes a semi-permeable membrane structure 115, asubstrate 110, a first microchamber 101 (hereinafter also referred to asthe sensing chamber) and a second microchamber 102 (hereinafter alsoreferred to as the reference chamber). Each of the two microchambers 101and 102 is formed between the semi-permeable membrane structure 115 andthe substrate 101, and includes a suspended element 120 and 130,respectively. The first microchamber 101 is adapted to receive asolution 150 including the polymer, and the second microchamber isadapted to receive a reference solution 160 for screening effects notcaused by the target analyte.

As shown in FIG. 1A, the two microchambers 101 and 102 are isolated fromeach other (not fluidically connected), each sealed by the substrate110, chamber side walls 103, 104, and 105, respectively. Theconfiguration of the two microchambers can be identical. Thesemi-permeable membrane structure 115 can be a continuous semi-permeablemembrane that covers and seals the two microchambers. Alternatively, thesemi-permeable membrane can include two semi-permeable membrane portionsthat each covers and seals the sensing chamber and reference chamber,respectively. The semi-permeable membrane structure is permeable to thetarget analyte 170 and impermeable to the polymer. Therefore, when thesample that can contain target analyte 170 is placed in contact with thesemi-permeable membrane structure, the target analyte 170, if present inthe sample, permeates through the semi-permeable membrane structure 115and enters the first microchamber 101 and the second microchamber 102,respectively.

The sensing chamber can be loaded with a solution 150 that includes apolymer, also referred to as the sensing polymer or targetanalyte-sensitive polymer hereinafter, that binds with the analyte. Thesensing polymer is prevented from escaping from the first microchamberthrough the semi-permeable membrane structure. The sensing polymer canbe biocompatible. In one embodiment, the biocompatible polymer canreversibly bind to the analyte of interest. The binding between thepolymer and the analyte can result in changes of the physicalcharacteristics (e.g., the viscosity and/or permittivity) of the polymersolution, which can be measured to determine the presence and amount ofthe analyte in the sample.

For example, when the analyte is glucose, through proper adjustment ofthe composition percentage of the boronic acid moieties on the polymerand polymer concentrations, the polymer can detect and differentiateglucose from other monosaccharides and disaccharides. Applying thispolymer to the sensor as disclosed herein can enable highly reliable,continuous monitoring of glucose in ISF in subcutaneous tissue.

As noted, the binding between the polymer and the analyte of interestcan be reversible. For example, the binding and dissociation between thetarget analyte and the sensing polymer can be an equilibrium phenomenondriven by the concentration of the analyte in the sensing chamber. Asthe analyte can move freely in and out of the sensing chamber throughthe semi-permeable membrane which the polymer cannot, the amount of theanalyte bound with the sensing polymer depends on the concentration ofthe analyte in the sample.

In one embodiment, a suitable polymer having boronic acid moieties canbe formed as a copolymer of at least two monomers, where one of themonomers includes at least one boronic acid functional group. Acopolymer can be synthesized with these monomers via classic freeradical copolymerization processes. In various embodiments, a suitablepolymer includes, but is not limited to, a polymer that contains boronicacid groups, or other receptor groups that recognize the given analytes.In one embodiment, the polymer is PAA-ran-PAAPBA, which is anamphiphilic copolymer containing two components, hydrophilic polymersegment polyacrylamide (PAA) and hydrophobic polymer segmentpoly(3-acrylamidophenylboronic acid) (PAAPBA).

A solution of PAA-ran-PAAPBA can undergo a viscosity change as well as apermittivity change when interacting with glucose molecules, asdiscussed in US Patent Application Publication No. 20120043203, assignedto the common assignee, the disclosure of which is incorporated hereinby reference in its entirety. In another embodiment, the sensing polymeris PHEAA-ran-PAAPBA, which is an amphiphilic copolymer containing twocomponents: PAAPBA and poly(N-hydroxyethyl acrylamide) (PHEAA). PAAPBAis a hydrophobic glucose-sensitive component, while PHEAA is ahydrophilic and nonionic component, and primarily serving to improve theoverall water solubility of the entire copolymer. When added to anaqueous solution of PHEAA-ran-PAAPBA, similar to PAA-ran-PAAPBA, glucosebinds reversibly to the phenylboronic acid moieties in the PAAPBAsegments to form strong cyclic boronate ester bonds, while having almostno response to other potential interferents, such as fructose,galactose, and sucrose.

To screen out effects not caused by the target analyte, for example,environmental factors such as temperature, the reference chamber canalso be loaded with a solution of another polymer (the referencepolymer). The reference polymer does not bind with the target analyte.Also, the reference polymer should not bind with or otherwise react withother substance in the sample solution to impact the property of thereference solution in a similar way as the target analyte impacts thecorresponding property in the sensing polymer solution. The referencepolymer can be selected to have similar hydrophilic blocks to those inthe sensing polymer, but have no phenylboronic acid moieties. Forexample, glucose-unresponsive PAA or PHEAA can be used as a referencepolymer for glucose detection. The viscosity of PAA (or PHEAA) solutionis glucose-independent. The analyte-free viscosity of the sensingpolymer solution can be similar to that of the reference polymersolution.

When the microdevice is viscosity-based, as shown in FIG. 1A, thesuspended elements 120 and 130 of the sensing chamber and the referencechamber can each act as a vibrational diaphragm, which can be actuatedby an external alternating field. The suspended elements (120, 130) eachcan include structural elements (121, 131), e.g., made from parylene,for structural integrity, passivation, and support for other components.For example, suspended element 120 and 130 can each include amagnetically active component 122 and 132, respectively, e.g., made of amagnetic material such as permalloy.

When an alternating electromagnetic field is applied, the suspendedelement 120 (as well as 130) can vibrate. The source of theelectromagnetic field, its relative configuration with the microdevice,and the mechanism in which the varying electromagnetic field interactwith the vibrational diaphragm can be as disclosed in US PatentApplication Publication No. 20120043203, or modifications thereof aswill be appreciated by those skilled in the art. Therefore, when thetarget analyte binds with the polymer, the viscosity change of thepolymer solution can influence the vibration of the suspended element120, while the vibration of the suspended element 130 in the referencechamber will not change. The vibration of the suspended element 120 canbe measured by the capacitance of a capacitor formed by a top electrode123 included in the suspended element 120, a bottom electrode 124 formedon the substrate, and an air gap 125 therebetween.

Similarly, the vibration of the suspended element 130 can be measured bythe capacitance of a capacitor formed by a top electrode 133 included inthe suspended element 130, a bottom electrode 134 formed on thesubstrate, and an air gap 135 therebetween. The top and bottom electrodecan be made from any common materials suitable for use in electrodes,such as gold, copper, other metals or alloys. By measuring thedifference of the vibrational behaviors, such as magnitude, of thesuspended element 120 and 130, the presence and/or amount of the analytein the sample can be determined.

FIG. 1B illustrates the structure of an example permittivity-basedmicrodevice. The reference numerals in FIG. 1B represent correspondingelements in FIG. 1A. While similar to FIG. 1A, the suspended element 120and 130 each include openings, or perforations, that allow the polymersolution and the reference solution to fill the gap between therespective suspended element and the substrate. In addition, thesuspended elements can also be supported by anti-stiction post (128 and138) formed from the substrate for structural stability. In this manner,the capacitance of the capacitor of formed between the top electrode123, bottom electrode 124, and the polymer solution filled therebetweencan be measured to detect any permittivity change caused by the analyte,and compared with the permittivity change of the reference solution. Insome embodiments, the device does not include a reference chamber, andthe permittivity change of the sensing chamber is directly correlatedwith the analyte presence and/or concentration.

Exemplary techniques for fabrication of the devices illustrated in FIGS.1A and 1B will be discussed in further detail in Example 1 and Example4.

Semi-permeable membranes are well known in the art. The semi-permeablemembranes suitable for the microdevices of the disclosed subject mattercan be obtained from commercial sources and selected based on pore sizesor cut-off molecular weight. In one embodiment, the semi-permeablemembrane is selected to be cellulose acetate.

In certain embodiments, the disclosed subject matter provides animplantable monitor comprising a MEMS affinity device as described abovecoupled with a wireless interface. The wireless interface can include acapacitance digital converter coupled with the microdevice and adaptedto produce a digital signal representing a measurement of the targetanalyte in the interstitial of the subject; a microcontroller coupledwith the capacitance digital converter; and a transponder coupled withthe microcontroller to transmit the digital signal received from thecapacitance digital converter to an external reader.

In various embodiments of the disclosed subject matter, the sensor canbe used to determine the level of an analyte in the body, for exampleoxygen, lactase, insulin, hormones, cholesterol, medicaments, viruses,or the like. The sensor can use any known method to provide an outputsignal indicative of the concentration of the target analyte. The outputsignal is typically a raw data stream that is used to provide a usefulvalue of the measured analyte concentration. In general, before thedevices is used to detect or monitor a target analyte, they are firstcalibrated using samples containing known amount of the target analyteto obtain correlations between sensor response (e.g., capacitancereadout) and the known concentration of the calibration sample.Thereafter, in the monitor of the target analyte, the pre-establishedcorrelations can be used to interpret the output signals of the sensorand determine the presence and/or concentration of the target analyte ina test sample.

In certain embodiments of the disclosed subject matter, the sensor isused to monitor glucose as the target analyte. In these embodiments, thesensor can measure a concentration of glucose or a substance indicativeof the concentration or presence of the glucose by using a specificpolymer in the sensor.

The sensor can also be used for other applications. In addition todiabetes, the proposed miniature CGM device can also be used for glucosemonitoring for other diseases (e.g., glycogen storage disease andhyperinsulinaemic hypoglycaemia).

The method can be extended to other metabolites, such as lactate, fattyacids, cysteines and homocysteines. For example, in emergency medicine,lactate monitoring can be used to predict possible organ failure oftrauma patients, organ transplant patients, and patients with othercritical conditions.

Further, the methods disclosed herein can be used as a reliable methodfor long-term monitoring of metabolites. Such methods can have greatmilitary significance. For example, a miniature device for glucosedetection with fully electronic readout would have significantapplications in protecting armed forces in the field. It can alsoprovide a platform to enable the delivery of drug treatments andnutritional supplements to protect and enhance performance in militarypersonnel.

Moreover, the disclosed method can be applied to the diagnosis ofdisease. For example, the development of boronic acid based glucosesensing systems can be extended to other analytes, such as human virusesand bacteria, since many of those microorganisms carry glycoproteins onthe exterior surface that can be targeted by the boronic acid basedbinding motifs.

Metabolic monitoring is of great utility to environmental monitoring.Changes in the concentrations of metabolites are the precursors andproducts of enzymatic activity, and can be associated with biologicalfunction and regulation. Metabolic monitoring hence can be used forenvironmental monitoring, e.g., risk assessment of chemicals anddiagnosis of diseases in wild animals. It can also be used as a tool tobetter understand the underlying mechanisms of action of toxic compoundsin the environment.

Additional aspects and embodiments of the disclosed subject matter areillustrated in the following examples, which are provided for betterunderstanding of the disclosed subject matter and not limitation. In allthe examples, glucose is used as the target analyte. As such, forconvenience, the disclosed microdevices will be also referred to asglucose sensors.

Example 1 A MEMS Differential Viscometer

Affinity glucose sensors based on viscosity detection using cantileveror diaphragm based vibrational elements are disclosed, e.g., in U.S.Patent Application Publication No. 20120043203. These devices caninclude a single sensing element (hence are also referred to as“single-module” devices herein) that measures glucose-induced viscositychanges in PAA-ran-PAAPBA polymer solutions. The results havedemonstrated the feasibility of these devices in CGM. However,requirements for closed-loop temperature control and minimumenvironmental disturbances during the sensor operation can presentdifficulties for implantable applications for these sensors, in whichthe simplicity in device design and stability in device performance aredesired.

Illustrated herein is a MEMS differential glucose sensor that can rejectundesired common mode interferences through differential measurements,allowing accurate glucose detection. This sensor includes twomagnetically driven vibrating diaphragms each situated inside amicrochamber. One of the microchambers (the sensing chamber) is filledwith a sensing solution of PHEAA-ran-PAAPBA, while the othermicrochamber (reference chamber) contains a reference solution of PHEAAthat does not bind or otherwise react with glucose or other componentsof the sample under analysis. As glucose permeates through asemi-permeable membrane into each chamber, the viscosity of the sensingpolymer solution increases due to glucose binding, while the viscosityof the reference solution only changes with environmental disturbances.Thus, measurement of the viscosity difference between the two chambersthrough differential capacitive detection of the vibration dampingallows determination of the glucose concentrations while rejectingcommon mode disturbances. In-vitro and in-vivo data as described inconnection with FIGS. 7, 8, 12-15 demonstrate the value of this sensorfor highly stable subcutaneous CGM applications.

The structure of an example CGM sensor is illustrated in FIG. 1 (sideview). FIG. 2 is a top view of the sensor as fabricated according to themethod disclosed in connection with FIG. 4. A pair of surface machinedfreestanding diaphragms (120, 130), one situated inside a sensingmicrochamber 101 (sensing diaphragm 120) while the other inside areference microchamber 102 (reference diaphragm 130), each vibrate underan external AC magnetic field. A top (moving) electrode (123, 133) isembedded in each of the diaphragms and is separated from a fixedelectrode (124, 134) on a substrate below by an air gap, forming adiaphragm position-sensing capacitor. Magnetically responsive permalloythin-film strips (122, 132) are integrated on each of the diaphragms(120, 130) and are passivated along with the moving electrodes to avoiddirect contact with the polymer solution 150 and reference solution 160.The CGM sensor as depicted in FIG. 2 also includes an inlet 181 andoutlet 182 for introducing and withdrawing the polymer solution, bondingpads 190 for providing electric connectivity for the electrodes, andetching hole 185 for etching the photoresist, as will be furtherdescribed in connection with FIG. 4.

The glucose sensitive polymer PHEAA-ran-PAAPBA utilized in the device isa synthetic polymer that recognizes glucose by specific affinitybinding. Specifically, PHEAA-ran-PAAPBA is an amphiphilic copolymercontaining two components: PAAPBA and poly(N-hydroxyethyl acrylamide)(PHEAA) (FIG. 3A). PAAPBA is a hydrophobic glucose-sensitive component,while PHEAA is a hydrophilic and nonionic component, and primarilyserving to improve the overall water solubility of the entire copolymer.When added to an aqueous solution of PHEAA-ran-PAAPBA, glucose bindsreversibly to the phenylboronic acid moieties in the PAAPBA segments toform strong cyclic boronate ester bonds, resulting in an increase in theviscosity of the solution (FIG. 3B), while having almost no response toother potential interferents, such as fructose, galactose, and sucrose.Glucose-unresponsive PAA is used as a reference polymer. The viscosityof PAA solution is glucose-independent.

To fabricate the device as shown in FIGS. 1A and 2, chrome (5 nm) andgold (100 nm) were first deposited and patterned to form two electrodes424 and 434 (500 μm×500 μm) on a SiO₂ coated silicon wafer 410 (FIG.4A), followed by the deposition of a parylene passivation layer 431 (1μm in thickness). A sacrificial photoresist layer 435 (5 μm) was thenspin-coated and patterned to define electrode air gaps (FIG. 4B),followed by the deposition of an additional parylene layer 437 (1 μm inthickness) (FIG. 4C). A second layer of chrome (5 nm) and gold (100 nm)were next deposited for the top electrodes 423 and 433 and Permalloyseed layer. Subsequently, defined by a photoresist pattern (5 μm inthickness), strips of Permalloy 422, 432 (220×35×3 μm) wereelectroplated (FIG. 4D). This was followed by removal of thephotoresist, patterning of the top electrodes (500×500 μm), anddeposition of an additional parylene passivating layer 447 (3 μm) and analuminum mask (458). Eight etching holes (250×250 μm) were openedthrough the parylene layers and reactive ion etching to expose thesacrificial photoresist (FIG. 4E), which was then removed by acetone(80° C.) to release the diaphragms (520×520 μm) (FIG. 4F). These etchingholes were then sealed by epoxy (Devcon Inc.). After wafer dicing (FIG.5) and wire bonding, a chip was bonded to an SU-8 sheet (thickness: 80μm), in which holes (1 mm in diameter) of appropriate sizes werepatterned to define the microchambers (0.06 μL in volume) as well as theinlets and outlets for the polymer solution handling. This SU-8 sheetwas in turn bonded using epoxy to a regenerated cellulose acetatesemi-permeable membrane (Membrane Filtration Products, Inc.) with asuitable thickness (for example 20 μm) and a suitable molecular weightcutoff (e.g., 6000 Da).

Chemicals and reagents used in the example include PHEAA-ran-PAAPBA andPAA, which were synthesized in house by free radical polymerization, asdescribed, for example, in Li S, et al., “Synthesis and development ofpoly(N-hydroxyethyl acrylamide)-ran-3-acrylamidophenylboronic acidpolymer fluid for potential application in affinity sensing of glucose.”J Diabetes Sci Technol 5:1060-1067 (2011). D-(+)-glucose were purchasedfrom Sigma-Aldrich. PBS, pH 7.4, was prepared by diluting a Ringer'sstock solution (Nasco Inc.) with sterile water (Fisher Scientific) at aratio of 1:9.

To prepare the sensing polymer solution, 284 mg of PHEAA-ran-PAAPBA withan hydroxyethylacrylamide (HEAA) to AAPBA molar ratio of 20 (orapproximately 5% of PAAPBA content in the polymer) and a molecule weightof 188,600, was dissolved in 6 mL of PBS, while the reference polymersolution was prepared by dissolving 142 mg of PAA in PBS (6 mL). Glucosestock solution (1 M) was prepared by dissolving glucose (180 mg) in PBSto 10 mL. A series of glucose concentrations (60, 90, 180, 360, and 500mg/dL) were prepared by further diluting the stock solution with PBS.

During testing, the microchambers were filled with solutions ofPHEAA-ran-PAAPBA and PAA, respectively. To facilitate in-vitro devicecharacterization, a test cell (volume: 300 μL) was constructed from anacrylic sheet directly above the MEMS sensor. A glucose solution at agiven concentration was introduced into the test cell, where it wasallowed to permeate through the semi-permeable membrane of the sensor tointeract with PHEAA-ran-PAAPBA in the sensing chamber. Because thevolume of the test cell was 5000 times larger than the microchambers, itwas reasonable to assume that the glucose concentration inside themicrochambers equalized to the given glucose concentration in the testcell when the glucose permeation reached an equilibrium.

FIG. 6 is an exemplary setup for characterization of the MEMSdifferential glucose sensor. The diaphragm vibrations were excited andmeasured using the setup in both the in-vitro and in-vivo tests. Acylindrical magnet 610 attached perpendicularly to the shaft 615 of abrushless DC motor 620 (Anaheim Automation) powered by power supply 630.The cylindrical magnet 610 can be spun with a maximum rated speed of4000 RPM. The motor shaft 615 is parallel to the plane of the MEMSsensor 100 and perpendicular to the Permalloy strips in the sensor,which were hence subjected to an AC magnetic field, inducing thevibration of the diaphragms. The average diaphragm vibration amplitudeswere measured by a capacitance digital converter (CDC) 640 which iscoupled with the MEMS sensor 100 and also communicates with a computer650. The CDC used herein was a Σ-Δ CDC (Analog Devices, AD7746), whichconverted the amount of charges on the sensor electrodes to acapacitance value. The CDC is capable of measuring a capacitance changeof ±4 pF with a measurement resolution and accuracy at 4 aF and 4 fF,respectively. To measure the differential capacitance, the CDC appliedan AC excitation voltage to the fixed electrodes of the sensing andreference diaphragms, while the moving electrodes in the diaphragms wereconnected to the C+ and C− pins of the CDC, respectively. Thecapacitances of the sensing and reference electrodes at rest were 81.5and 63.5 pF, respectively, both of which were beyond the measurementrange of the CDC. As a result, the nominal excitation voltage of the CDC(3.3 V) was trimmed to accommodate the measurement range (±4 pF).

The device temperature was uncontrolled throughout the tests, except forthe characterization of device temperature stability, in which thedevice was varied among physiological temperature via closed-looptemperature control. In in-vivo testing, the sensor was implanted in thesubcutaneous tissue of a sedated mouse whose glucose concentration wascontrolled by glucose and insulin injections. The implanted glucosesensor continuously measured the glucose level in ISF, while acommercial glucometer (Freestyle Lite®) sampled blood sugar levels inthe mouse's tail at specified frequencies.

Sensor Response to Glucose at Physiological Relevant Concentrations

Sensor response to various glucose concentrations was investigated tocharacterize the device resolution in glucose detection. Thedifferential diaphragm vibration was measured at physiologicallyrelevant glucose concentrations while the magnet spun at a fixedfrequency of 13 Hz. As the glucose concentration changed from 0 to 500mg/dL, the differential capacitance of the device decreased steadilyfrom 28 to 27.4 pF, reflecting a decrease in the difference of thediaphragm vibration amplitudes and an increase in the viscous dampingdue to glucose binding with PHEAA-ran-PAAPBA (FIG. 7). The deviceresolution is, thus, determined to be approximately 0.003 mg/dL with ameasuring accuracy of 3.3 mg/dL, which is sufficient for implantableglucose sensors. In addition, the results exhibit increased timeconstants at higher glucose concentrations. This can be caused by theincrease in the viscosity of the polymer solution, which slowed down theglucose diffusion. Sensor capacitance was also observed to becomesaturated at high glucose concentrations. As a result the relationbetween glucose concentrations and the differential capacitance isnonlinear and can be represented by a quadratic equation. As thesensitivity of glucose sensors typically changes after implantation, therelation obtained in-vitro is only qualitative for in vivo sensorcalibration. The glucose response of the device has demonstrated thatthe device can detect physiologically relevant glucose concentrationsand can be used in implantable glucose monitoring.

The sensor was also exposed to glucose solutions whose concentrationswere changed back and forth between two different values to characterizethe device time responses. For example, glucose concentration wasinitially allowed to be equilibrated at 60 mg/dL in the test cell andmicrochambers. Next, the solution in the test cell was replaced withanother glucose solution at 90 mg/dL. When the glucose concentrationinside the microchambers had equilibrated to 90 mg/dL, the reverseprocess was initiated, in which the test cell was refilled with a 60mg/dL glucose concentration. The process of solution refilling of thetest cell (within 10 seconds) was sufficiently fast as compared to theglucose concentration equilibration. During the equilibration processes,the differential sensor capacitance, at a fixed frequency of 13 Hz, wasmeasured as a function of time.

The result of the time response measurement is shown in FIG. 8. As theglucose concentration varies from 60 to 90 mg/dL, the differentialsensor capacitance decreases with time, corresponding to a decrease inthe sensing diaphragm vibration amplitude as well as an increase in theviscous damping to the diaphragm vibration. The capacitance finallysaturates to a constant level, reflecting that the process of glucosepermeation and binding have reached a dynamic equilibrium. Consideringthe glucose concentration change as a step input, the time constant ofsystem's step response is determined to be approximately 1.48 minutes.In the reverse process, the glucose concentration in the test cell isdecreased from 90 to 60 mg/dL. The sensor capacitance increases withtime, indicating an increase in the vibration amplitude of the sensingdiaphragm due to the reduced viscous damping. The time constant for thereverse process is approximately 2 minutes. The longer reverse timeconstant could be due to the smaller diffusivity of glucose molecules inthe initially more viscous polymer solution. Time constants at otherglucose concentrations can be obtained from FIG. 7, in which an averagetime constant of approximately 3 minutes is observed. Note that thesetime constants compare favorably with response times of commerciallyavailable systems, which range from 5 to 15 minutes.

The reversibility of the device response can be obtained by comparingdifferences in sensor output between two separated measurements at thesame glucose concentration. For example, as shown in FIG. 8, the sensoroutput varies from 27.642 (averaged between 0 and 7 minutes) to 27.485pF (averaged between 12 and 27 minutes) as the glucose concentrationvaries from 60 to 90 mg/dL. The sensor output then returns to 27.636 pF(averaged between 32 and 36 minutes) when the glucose concentration isreversed to 60 mg/dL. The difference between the average sensor outputsover the two periods with the glucose concentration at 60 mg/dL is onlyabout 6 fF, or 217 ppm. Note that this reversibility is achieved withouttemperature control, and is acceptable for implantable applications.

The time constant of the device can be effectively represented byglucose diffusion time in the sensing microchamber, which can beassessed by a simulation using COMSOL Multiphysics with a simplifieddevice model as depicted in FIG. 9 at different initial glucose levelsand various glucose concentration changes. As shown in FIG. 9, adiaphragm sealed inside a microchamber by a semi-permeable membrane hasbeen considered as a two-dimensional rectangular region, which containsfour walls labeled from 1 to 4. Among them, wall 1 and wall 4 separatedby a distance of 80 μm represent the membrane and the sensor diaphragmrespectively. And wall 2 and wall 3 are considered sidewalls of themicrochamber separated by a distance of 1 mm. The resultant arearepresents the microchamber filled with a PHEAA-ran-PAAPBA polymersolution at various glucose concentrations. The glucose diffusioncoefficient in water is selected to be 7×10⁻¹⁰ m²/s which can be used todetermine other constants in the Einstein-Stokes equation:D_(g)=k_(B)T/(6πηr), where D_(g) is the glucose diffusion coefficient.k_(B) is the Boltzmann's constant. T is the absolute temperature. η isthe viscosity of the polymer solution and r is the radius of the glucosemolecule.

Simulated device time responses was first obtained when the glucoseconcentration changes from 60 to 90 mg/dL and then reserves back to 60mg/dL. In the simulation, the glucose concentration at wall 1 is fixedat 90 mg/dL, while the rectangular region has an initial glucoseconcentration of 60 mg/dL. In the reverse process, the glucoseconcentration at wall 1 is fixed at 60 mg/dL, while the initial glucoseconcentration in the rectangular region is 90 mg/dL. The time-dependentglucose concentration on wall 4 was obtained as shown in FIGS. 10 and11. The time constants for the glucose increase and decrease processesare 1.67 and 1.99 minutes, respectively. The simulation resultscorrectly predict the order of magnitude of the time constantsdetermined from the test data. In particular, these results indicatethat a large time constant in the reverse process, which is consistentwith these measurements.

The same simplified device model also allows for simulation of thedevice time responses to other glucose concentration changes, as havebeen presented in FIG. 7. The glucose concentrations in the sensingmicrochamber are initially at 0, 60, 180, and 360 mg/dL, respectively,and then gradually approach the glucose concentrations at wall 1 thatare correspondingly fixed at 60, 180, 360, and 500 mg/dL, respectively.The time constants of these processes are determined by simulation to be1.3, 1.7, 2.3, and 3.3 minutes, respectively, which are consistent withmeasured results (FIG. 7).

The ability of the device to resist the temperature variations was alsocharacterized. The temperature of the device was altered from 34 to 40°C. under a closed-loop controlled heating system at a fixed glucoseconcentration of 60 mg/dL. Both the differential capacitance andsingle-module capacitance were obtained (FIG. 12). As the devicetemperature changed by 6° C., the differential and the single-modulecapacitance changed by 0.8 and 2.72 pF, respectively, indicating thatthe differential measurements effectively compensated the interferencefrom the temperature variations. The compensation effect of differentialmeasurements to temperature variations can be only partial, as thesensing and the reference diaphragm can respond to temperature changesdifferently. However, in reality, much smaller and slower temperaturevariations can be expected in in-vivo applications.

The glucose-independent drift in differential and single-module sensorcapacitance was assessed at a fixed glucose concentration withoutcontrolling the temperature. The sensor was exposed to a 60 mg/dLglucose solution over an extended period about 5 hours (FIG. 13). Duringthis period, the CDC was programmed to record the differential and thesingle-module capacitance alternatively under the same environmentalconditions (e.g., temperature variations, lighting, and osmoticpressure). It was observed that the drift in differential sensorcapacitance was significantly smaller than the single-modulecapacitance. The differential capacitance of the sensor changes from27.24 to 27.23 pF (measured at 0 and 320 minutes respectively),indicating a drift at approximately 1.8 fF/hour, while this valuebecomes 94 fF/hour in the single-module measurements. The large drift inthe single-module measurements is probably caused by the temperaturevariations and osmotic pressure, which are largely compensated in thedifferential measurements. These results demonstrate that thisdifferential device is capable of compensating for environmentaldisturbances and providing excellent stability, which is ideal forlong-term implantable CGM.

In-vivo characterization of the device was performed with a laboratorymouse. The glucose sensor was implanted in the subcutaneously tissue ofa sedated lab mouse to measure the glucose concentration in ISFcontinuously (FIG. 14A), while a commercial glucometer sampled theglucose level in the capillary blood from the tail tip of the mouseevery 10 minutes after glucose injection and every 5 minutes afterinsulin intervention. FIG. 14B shows the measurement data, which aregiven in terms of the change of the differential capacitance, calculatedwith respect to the value at the time of the first glucometer reading.It can be seen that this device output closely follows the commercialglucometer readings as the mouse's blood sugar levels vary over a 3-hourperiod. This result indicates consistency between the sensor output andthe glucometer reading, and supports use of this glucose sensor forlong-term CGM.

The differential sensor capacitance (C_(out)) was calibrated to obtainestimated blood glucose values (Ĝ₁). A quadratic equation was used torepresent the relation between C_(out) and ISF glucose concentrations(G₂). Here, G₂ can be expressed by C_(out) asG ₂ =aC _(out) ² +bC _(out) +c  (1)where a, b, and c are constants that can be determined from G₂ andC_(out). Due to the mass transfer of glucose, a physiological time lagof the concentrations between the ISF glucose and the blood glucoseexists. The kinetics of glucose concentrations in blood (G₁) and ISF(G₂) can be expressed asdG ₂ /dt=−(k ₀₂ +k ₁₂)G ₂ +k ₂₁ V ₁ /V ₂ G ₁  (2)where k₁₂ is the flux rate for forward glucose transport acrosscapillaries and k₂₁ is the flux rate for reverse glucose transportacross capillaries. k₀₂ is the glucose uptake into subcutaneous tissues.V₁ and V₂ are volumes of the blood and the ISF respectively. k₁₂, k₂₁,k₀₂, V₁, and V₂ are all constants. The combination of equation (1) and(2) yieldsG ₁ =a ₁ C _(out) ² +a ₂ C _(out) +a ₃ dC _(out) ² /dt+a ₄ dC _(out)/dt+a,  (3)where a₁, a₂, a₃, a₄, and a₅ are constants that can be determined bypartial least squares fitting using six blood glucose values (G₁) fromglucometer and their corresponding C_(out) from the implanted sensor.After the determination of a₁, a₂, a₃, a₄, and a₅, the estimated glucosevalue G₁ can be obtained using equation (3) with the known a₁, a₂, a₃,a₄, a₅, and C_(out).

The clinical accuracy of the Ĝ₁ as compared to G₁ can be quantifiedusing a Clarke error grid, which has been divided into several zones(e.g., A, B, C, D, and E) to represent different levels of accuracy. Forexample, if a point falls into Zone A or Zone B, the measurement iseither clinically accurate or clinically acceptable. In contrast, if apoint falls into another zone, then that the measurement might lead toproblems, such as overcorrection, dangerous failure, and erroneousness.In these measurements, all points in the Clarke error grid exclusivelyfall into Zone A (95.3%) and Zone B (4.7%), while no point falls intoother zones (FIG. 15). These results indicate a good clinical accuracyof the measurements, showing a great promise to apply this affinitysensor for long-term in-vivo glucose monitoring.

The results as described in connection with FIGS. 7-15 in this Exampleestablish that the sensor experienced a decrease in the differentialcapacitance of approximately 0.6 pF when the glucose concentrationincreased from 0 to 500 mg/dL. The time constant of the sensor wasapproximately 1.48 minutes during a glucose concentration change from 60to 90 mg/dL. The sensor also exhibited excellent reversibility; thedifferential capacitance of the sensor measured at two separatedmeasurements at 60 mg/dL glucose concentration agreed within 99.97%.

In addition, by varying the device temperature from 34 to 40° C. at aglucose concentration of 60 mg/dL, the differential capacitance changedby 0.8 pF, which is at least three times smaller than the change in thesingle-module capacitance, indicating the sensor's ability in resistingtemperature variations. By exposing the sensor to a 60 mg/dL glucoseconcentration for an extended measurement period of 5 hours, thedifferential sensor output exhibited low drift (1.8 fF/h), which isappropriate for long-term, stable CGM. Moreover, the results asdescribed in connection with FIGS. 14-15 indicate that the sensor outputclosely follows blood glucose concentrations in laboratory mice asmeasured by a commercial glucometer. Clarke error grid analysis usingthe calibrated in-vivo sensor data (FIG. 15) have demonstrated theclinically accuracy of the sensor measurements.

Example 2 A Dielectric Affinity Sensor

U.S. Patent Application Publication No. 20120043203 discloses adielectric CGM sensor that has no moving structure, and thus can bestable in face of environmental disturbances. In this example, adielectric glucose sensor with a perforated electrode is described. Thechange in permittivity in the polymer solution as a result ofglucose-polymer binding can be measured from the capacitance of thecapacitor formed between the electrodes. Results from in-vitrocharacterization of this sensor as described in FIGS. 20-26 demonstratethat this dielectric sensor can be useful in CGM.

FIG. 16 shows a schematic diagram of a single-module MEMS dielectricglucose sensor used in this Example. The sensor includes a microchamber1601 filled with a glucose-sensitive polymer solution 1650 and sealed bya semi-permeable membrane 1615. A perforated electrode 1623 embedded ina diaphragm 1630 is separated from a bottom electrode 1624 (which alsocan be perforated) on a substrate 1610 below by the polymer solution1650. Environmental glucose 1670 that permeates through thesemi-permeable membrane binds with the polymer and changes thepermittivity of the polymer solution. As the polymer solution canpermeate through the top perforated electrode 1624, it can fill the gapbetween the perforated electrode 1624 and substrate 1610. Theanti-stiction posts 1638 can support the diaphragm 1630 and prevent thediaphragm from collapsing while providing additional resistance toenvironmental disturbances (e.g., shock, vibration, etc.)

The perforated electrode 1623 embedded in the diaphragm 1630 forms aparallel plate capacitor with the passivated bottom electrode 1624 onthe substrate. The glucose-sensitive polymer used here wasPAA-ran-PAAPBA, which interacts with glucose by specific affinitybinding as described in U.S. Patent Application Publication No.20120043203. In brief, when added to an aqueous solution ofPAA-ran-PAAPBA, glucose binds reversibly to phenylboronic acid moietiesin AAPBA segments to form strong cyclic boronate ester bonds, resultingin a change in the permittivity of the dielectric solution as well as achange in the sensor capacitance.

In an electric field, a number of polarization mechanisms contribute tothe permittivity of the polymer solution as well as the measured sensorcapacitance. These polarization effects are time-dependent in a harmonicelectric field, and are influenced by the polymer molecular structure.The solution of PAA-ran-PAAPBA can undergo a molecular structure changewhen the polymer binds to the glucose. Thus, at a given frequency, thepermittivity can change, which can be measured to determine the glucoseconcentration. The permittivity of the polymer solution is a complex andcan be written as ∈*=∈′−i∈″, in which the capacitive component ∈′represents the ability of the polymer solution to store the energy fromthe electric field, while the resistive component ∈″ is related toenergy loss. As ∈″ is directly proportional to the device capacitance(C_(x)), any changes in ∈′ can be determined from capacitancemeasurements.

The fabrication of the device started with deposition and patterning ofa thin film gold layer to form a bottom electrode (1 mm×1 mm×100 nm) aswell as a resistive temperature sensor on a SiO₂ 1711 coated siliconsubstrate 1710 (FIG. 17A). A parylene passivating layer 1731 (1 μm inthickness) was then deposited by chemical vapor deposition. Followingthe spin-coating and patterning of a sacrificial photoresist layer 1735(3 μm in thickness) (FIG. 17B), an additional parylene layer 1737 (1.5μm in thickness) was deposited. A gold layer was further deposited andpatterned to form a perforated electrode 1733 (FIG. 17C), which was thenpassivated by another parylene layer 1738 (3 μm in thickness) (FIG. 17D)and a patterned SU-8 reinforcement layer 1739 (20 μm in thickness) (FIG.17E), resulting in nine anti-stiction posts 1748 with diameters of 50μm. A SU-8 layer 80 μm in thickness was finally spin-coated andpatterned to form microchamber wall 1733 as well as an inlet and anoutlet for polymer solution handling. The two successively coated SU-8layers also acted as a mask for patterning of the underneath parylenelayers by reactive ion etching to expose the sacrificial photoresistlayer, resulting in a diaphragm with holes for glucose diffusion. Thediaphragm was at last released by removal of the sacrificial layer in aphotoresist stripper. A cellulose acetate semi-permeable membrane 1715(Membrane Filtration Products, Inc) was in turn glued onto themicrochamber 1701 (FIG. 17F) by epoxy (Decvon Inc). The sensor wasencapsulated into an acrylic test cell with a total volume ofapproximately 1 mL. FIG. 18 shows images of the sensor before packaging.

The PAA-ran-PAAPBA polymer was synthesized in house by free radicalpolymerization (see S. Li et al., “Development of Novel Glucose SensingFluids with Potential Application to MicroelectromechanicalSystems-Based Continuous Glucose Monitoring,” Journal of DiabetesScience and Technology, 2: 1066-1074, (2008); S. Li, et al.,“Development of Boronic Acid Grafted Random Copolymer Sensing Fluid forContinuous Glucose Monitoring,” Biomacromolecules, 10: 113-118, (2008)).To prepare the polymer solution, 284 mg of PAA-ran-PAAPBA, with an AA toAAPBA molar ratio of 20 (or approximately 5% PAAPBA content in thepolymer) and a molecule weight of 170,700, was dissolved in 6 mL ofphosphate buffer saline (PBS). The PBS buffer, pH 7.4, was prepared bydiluting a Ringer's stock solution (Nasco) with sterile water (FisherScientific) at a ratio of 1:9. D-(+)-glucose was purchased fromSigma-Aldrich. Glucose stock solution (1 M) was prepared by dissolvingglucose (1.8 g) in PBS to 10 mL. A series of glucose solutions (30mg/dL, 60 mg/dL, 90 mg/dL, 120 mg/dL, 240 mg/dL, and 480 mg/dL) wereprepared by further diluting the stock solution with PBS.

The microsensor was characterized using the setup shown in FIG. 19. InFIG. 19, the MEMS sensor 100 was integrated into a capacitance/voltagetransformation circuit driven by a sinusoidal input from a functiongenerator 1910 (Agilent, 33220A). The sensor is also coupled with amultimeter 1920 for outputting measurement result. The temperature ofthe polymer solution in the MEMS sensor 100 was maintained at 37° C. viaclosed-loop control by a Peltier heater 1970 (Melcor, CP14) powered bypower supply 1930. The voltage of the Peltier heater 1970 is alsocontrolled according to the feedback from the on-chip temperaturesensor. Further details of FIG. 19 can be found in X. Huang, et al., “ADielectric Affinity Microbiosensor,” Appl. Phys. Lett., 96:033701-033703, (2010). All tests were conducted at frequencies below 100kHz as allowed by a lock-in amplifier 1950 (Stanford Research Systems,SR830), which measured the amplitude and the phase shift of the outputvoltage from the circuit, and communicates with the computer 1940. Theequivalent capacitance (C_(x)) that is directly related to the polymerpermittivity was determined from the circuit outputs when the MEMSsensor 100 and a reference capacitance 1980 (C_(R)) are in turn coupledinto the circuit by switching T between position S and R.

The device glucose response was measured under an E-field at a range ofdriving frequencies. The device's equivalent capacitance as a functionof frequency for the glucose-free PAA-ran-PAAPBA polymer solution wasfirst obtained. As shown in FIG. 20, the sensor capacitance decreasesconsistently with the frequencies due to the frequency-dependentdielectric relaxation of the polymer. In addition, a rapid decrease ofsensor capacitance from 72.7 to 20 pF was also observed with thefrequency changed from 5 to 20 kHz. This can be attributed tointerfacial polarization, which typically dominates at low frequency. Byexposing the device to various glucose concentrations ranging from 30 to480 mg/dL, the sensor capacitance decreases with increasing glucoseconcentrations at all measured frequencies (FIG. 21). For example, at100 kHz frequency, the sensor capacitance decreases by 0.3 pF at 480mg/dL with respected to the sensor capacitance in the glucose-freepolymer solution. These results suggest that the glucose concentrationcan be determined through permittivity measurement at a fixed frequency(e.g., 100 kHz).

The glucose-dependent permittivity or capacitance changes of the devicecan be attributed to a number of polarization mechanisms, such aselectronic polarization, ionic polarization, dipolar reorientation,counterion polarization, and interfacial polarization. First, theelectronic polarization and the ionic polarization are referred to thedistortion of electron cloud and displacement of ions in the appliedE-field, respectively. Second, the dipolar reorientation involvesalignment with the applied E-field of permanent dipoles, which, forPAA-ran-PAAPBA, can include AAPBA and AA segments rigidly attached tothe polymer backbone. Third, in the counterion polarization, appendinggroups of PAA-ran-PAAPBA are negatively charged, and cations (e.g., Na⁺,K⁺, and H₃O⁺) are attracted to form a counterion cloud. Under theE-field, the counterions migrate unevenly within the cloud to contributea net dipole moment.

The interfacial polarization involves dipole moments due to electricaldouble layers formed at the interfaces of the ionic buffer with polymermolecules (i.e., Maxwell-Wagner-Sillars polarization) as well as thepassivated electrode surfaces (electrode polarization). Theseinterfacial polarization effects dominate the low-frequency regions, andare generally exhibited as a sharp decline of permittivity withincreasing frequencies. The relaxation frequency of electronic and ionicpolarization is on the order of 1 THz, and interfacial polarization ison the order of 1 GHz, while those of dipole reorientation andcounterion polarization are on the order of a few kHz to a few tens ofkHz. Thus, all of these polarization mechanisms can be significant forthe polymer, and the relaxation behavior apparent from the rapid drop ofthe measured capacitance at frequencies lower than 20 kHz (FIG. 20) canbe mainly due to the interfacial polarization.

At the above measurement frequencies, the polarization behavior ofPAA-ran-PAAPBA is influenced by glucose binding. As AAPBA segments bindwith glucose at a two to one ratio to form cyclic esters of boronic acidby eliminating two hydroxyl groups. This can cause a decrease of netpermanent dipole moments, thereby reducing the energy storage ability ofthe polymer solution as well as the capacitive component (∈′) in thepermittivity. Furthermore, glucose binding can lead to variations in thenet charge of polymer segments as well as changes in the polymerconformations, which would alter the electric double layer structure andresult in changes in Maxwell-Wagner-Sillars and counterion polarization.Moreover, the crosslinking of polymer after glucose binding can increasethe elastic resistance of the permanent dipoles in polymer to alignmentwith the E-field, leading to a decrease in ∈′. The combination of theseeffects explains that at a given frequency, the measured sensorcapacitance decreased with glucose concentrations (FIG. 21).

To characterize the device time response, the glucose concentration wasallowed to be equilibrated at 60 mg/dL in the test cell and themicrochamber. Next, the solution in the test cell was replaced withanother glucose solution at 120 mg/dL. When the glucose concentrationinside the sensor chamber had equilibrated to 120 mg/dL, the reverseprocess was initiated, in which the test cell was refilled with aglucose solution at 60 mg/dL again. The process of glucose sampleintroduction was typically within a few seconds, which was sufficientlyfast when compared with the time for the glucose concentrationequilibration. Throughout this concentration equilibrium process, an ACvoltage of a fixed frequency of 100 kHz was applied to the sensor, andthe sensor capacitance changes at this frequency were obtained.

From the data (FIG. 22), it can be seen that as the glucoseconcentration varies from 60 to 120 mg/dL, the sensor capacitancedecreases with time, corresponding to a decrease in the permittivity ofthe polymer solution due to glucose binding. The sensor capacitancefinally saturates to a constant level, reflecting that the process ofglucose permeation and binding have reached a dynamic equilibrium.Assume that the glucose concentration change as a step input, the timeconstant of the device represents the time it takes the system's stepresponse to reach 63.2% of its final value. The time constants for theforward and reverse processes are approximately 2.49 and 3.08 minutesrespectively. The longer reverse time constant could be due to thesmaller diffusivity of glucose molecules in the initially more viscouspolymer solution and have been confirmed by the simulation in Example 1.

The reversibility of the device response can be obtained by comparingdifferences in sensor output between two separated measurements at thesame glucose concentration. For example, as shown in FIG. 22, the sensorcapacitance at 60 mg/dL glucose concentration varies from 11.951(averaged over the period between 0 and 5 minutes) to 11.952 pF(averaged over the period between 26 and 31 minutes). The differencebetween the average sensor outputs over the two periods with the glucoseconcentration at 60 mg/dL is only about 1 fF, indicating that the sensorpossesses excellent reversibility with respect to glucose concentrationvariations.

The drift of the device was investigated by exposing it to a constantglucose concentration (60 mg/dL) over an extended measurement period.The sensor capacitance at 100 kHz is shown in FIG. 23. It can be seenthat the sensor capacitance is steady at 11.955 pF over a period ofabout 4 hours with slight drift. The low drift demonstrates that thedevice can offer highly stable measurements for long-term continuousglucose monitoring. However, some fluctuations during the measurementwere also observed, which can be explained as the environmentaldisturbances, such as shocks, vibrations, and human activities, whichrandomly appear in the testing environment.

Example 3 Another Dielectric Affinity Sensor

In this example, the affinity sensor discussed in Example 2 is used,except that the polymer in the sensing chamber was changed topoly(N-hydroxyethyl acrylamide)-ran-3-acrylamidophenylboronic acid(PHEAA-ran-PAAPBA). Similar studies were performed on the frequencydependence of the sensor response, time response and drift of thesensor. FIG. 24 shows that in the absence of glucose, the sensorcapacitance decreased monotonically from 0.5 to 20 kHz, and thenincreased slowly at higher frequencies where orientational polarizationwas significant. FIG. 25 shows the sensor's time response assessed at afixed frequency of 100 kHz. In response to a step glucose concentrationchange from 50 to 100 mg/dL, the sensor showed a time constant of 4.7minutes, which is acceptable for CGM and further improvable byoptimizing the sensor geometry. The drift of the sensor over time isvery small, as shown in FIG. 26.

Example 4 A Differential Dielectric Affinity Sensor

In Examples 2-3, glucose sensors based on permittivity detection arediscussed. Each of these devices contains a single sensing element thatmeasures the glucose-induced permittivity changes in the polymersolutions. The results demonstrate the use of these devices in long-termand stable CGM. However, as the dielectric property of polymer solutionsis very sensitive to disturbances, these dielectric sensors requireclosed-loop temperature control to maintain the device temperature, andexhibit limited resistance to environmental interferences. As a result,noticeable fluctuations of the sensor signal can be observed in thedrift measurement (FIG. 23).

Differential sensing has been successfully applied in the development ofthe viscometric glucose sensor, as illustrated in Example 1, which showsan improved stability in face of common mode disturbances. In thisexample, a MEMS differential dielectric sensor having two microchambersutilizing permittivity measurement is discussed. The glucoseconcentration can be determined from the permittivity difference betweenthe sensing and the reference solutions, which is measured as thedifferential capacitance. The test results in this example as describedin connection with FIGS. 29-37 demonstrate that the sensor allowssensitive and specific detection of glucose at physiologically relevantconcentrations with improved stability to external interferences,showing a great promise to apply the differential dielectric sensor forfully implantable, long-term CGM.

The structure of this sensor is depicted in FIG. 1B. In brief, thissensor includes a pair of perforated electrodes (123, 133) situatedinside a sensing microchamber 101 and a reference microchamber 102,respectively. The sensing chamber 101 contains a solution of a glucosesensitive polymer 160, while the reference chamber 102 is filled with asolution of a reference polymer 160 that is not responsive to glucose.The microchambers 101 and 102 are sealed with a semi-permeable membrane,which prevents the polymers from escaping from the microchambers, whileallowing environmental glucose to diffuse through. The perforatedelectrodes (123, 133) are embedded in suspended diaphragms (120, 130)that are supported by arrays of anti-stiction posts (128, 138). Theseposts prevent the diaphragms from collapsing, while offering additionalsupport for the diaphragms from environmental disturbances. Each of theperforated electrodes is separated from a bottom electrode (124, 134) bya gap that is also filled with the sensing polymer solution 150 or thereference polymer solution 160, resulting in a capacitor with thepolymer solution as dielectrics. As glucose permeates through thesemi-permeable membrane into each chamber, the permittivity of thesensing polymer solution is changed due to glucose binding, while thepermittivity of the reference solution is unchanged due to a lack ofglucose binding. The permittivity difference between the sensing and thereference solution can be determined from differential capacitance,which also allows determination of the glucose concentrations whilerejecting permittivity changes caused by environmental fluctuations. Thedifferential sensor uses PHEAA-ran-PAAPBA as glucose sensitive polymer,and PAA as a reference polymer, same as described in Example 1.

The procedure to fabricate the differential sensor in this example isschematically shown in FIGS. 27A-27F, which are similar to FIGS. 17A-17F(which are for single-module sensor). A thin film gold layer wasdeposited and patterned to form bottom electrodes 2724 (1 mm×1 mm×100nm) as well as resistive temperature sensors on a silicon substrate 2710coated with silicon oxide 2711 (FIG. 27A). A parylene passivating layer2731 (1 μm in thickness) was then deposited by chemical vapordeposition. Following the spin-coating and patterning of a 5 μmsacrificial photoresist layer 2735 (FIG. 27B), an additional parylenelayer 2737 (1.5 μm in thickness) was deposited. A gold layer was furtherdeposited and patterned to form top perforated electrodes 2733 (FIG.27C), which was then passivated by another parylene layer 2738 (3 μm inthickness) (FIG. 27D) and a succeeding SU-8 reinforcement layer 2739 (20μm in thickness) (FIG. 27E), resulting in nine anti-stiction posts 2748with diameters of 50 μm for each perforated electrode. A SU-8 layer 80μm in thickness was finally spin-coated and patterned to formmicrochamber walls as well as inlets and outlets for polymer solutionhandling. The two successively coated SU-8 layers also acted as a maskto pattern the underneath parylene layers by reactive ion etching toexpose the sacrificial photoresist layer and form diaphragms with holesfor glucose diffusion. The diaphragms were released by the removal ofsacrificial layer in photoresist stripper. A CA semi-permeable membrane2715 (Membrane Filtration Products, Inc) was in turn glued onto the twomicrochambers 2701 and 2702 by epoxy (Decvon Inc) (FIG. 27F). Thedifferential sensor was encapsulated into an acrylic chamber with atotal volume of approximately 1 mL. FIG. 28 shows images of the sensorbefore and after packaging.

Two setups were used to characterize the device. First, acapacitance/voltage converter circuit similar to what is depicted inFIG. 19 was used to measure the frequency response of both the sensingand the reference polymer. Specifically, this circuit measures thesingle-module capacitance of the sensing or the reference electrodesunder an AC E-filed with an amplitude of 10 mV and a frequency variedfrom 0.5 to 100 kHz. The amplitude and the phase of the output voltagefrom the circuit are captured by a lock-in-amplifier (SR830) tocalculate the sensor capacitance using previously reported equations.This setup is further simplified using an Σ-Δ CDC (Analog Devices,AD7746), which converts the amount of charges on the capacitive sensorelectrodes to a capacitance value. This CDC processes a measurementcapacity of ±4 pF in capacitance changes with a resolution at 4 aF andan accuracy at 4 fF. To measure the differential capacitance, the CDCapplies a square wave at a frequency of 32 kHz to the bottom electrodes,while the perforated electrodes are connected to capacitance measurementpins of the CDC. The initial capacitances of the sensing and thereference electrodes were determined to be 57.8 and 19.2 pF,respectively, which were beyond the measurement range of CDC. Thus theeffective excitation voltage from the CDC was trimmed from a designatedvalue of 3.3 V to an appropriate value to accommodate the measurementrange. The CDC can be programmed to obtain either the capacitancedifference of the sensing and the reference electrodes or thesingle-module capacitance solely from the sensing electrode.

Chemicals and reagents used in the this example includePHEAA-ran-PAAPBA, which was synthesized in house by free radicalpolymerization with an HEAA to AAPBA molar ratio of 20 (or approximately5% PAAPBA content in the polymer) and a molecule weight of 188600. PAAwas also synthesized by a similar process as PHEAA-ran-PAAPBA usingacrylamide monomer. D-(+)-glucose was purchased from Sigma-Aldrich. PBS,pH 7.4, was prepared by diluting a Ringer's stock solution (Nasco) withsterile water (Fisher Scientific) at a ratio of 1:9. PHEAA-ran-PAAPBA(284 mg) and PAA (142 mg) were dissolved in PBS (6 mL) to obtain asensing and a reference solution, respectively. Glucose stock solution(1 M) was prepared by dissolving glucose (180 mg) in PBS to 10 mL. Aseries of glucose concentrations (50, 100, 200, 300, 400, and 500 mg/dL)were prepared by further diluting the stock solution with PBS.

Both in-vitro and in-vivo testing was performed to characterize thedevice. First, the frequency responses of the electrodes withglucose-free sensing and reference solutions were characterized usingthe capacitance/voltage converter circuit. In addition, the frequencyresponses of the sensing polymer at selected glucose concentrations werealso obtained. The time responses of the sensor upon glucoseconcentration changes were then obtained using the CDC, which was alsoused for all the following tests.

The frequency responses of the sensing and the reference electrodes tothe E-field with a frequency from 0.5 to 100 kHz were obtained using thecapacitance/voltage converter circuit. As shown in FIG. 30, when thepolymer solutions contained no glucose, the frequency responses of thesensing and the reference electrodes, which are represented as thecapacitances, decrease consistently from 0.5 to 20 kHz and afterwardsundergo slow increases. The abnormally decreases of the electrodecapacitances at low frequencies could be attributed to effects ofelectrode polarization and Maxwell-Wagner-Sillars polarization, whichtypically happen between the interface of two different media and areexhibited as a rapid decrease of permittivity at low frequencies.

The capacitance changes in the sensing electrodes after the bindingbetween glucose and the polymer were also obtained at selected glucoseconcentrations from 50 to 200 mg/dL. As shown in FIG. 30, the electrodecapacitance decreases with increasing glucose concentrations after 10kHz. In contrast, at frequencies between 0.5 and 10 kHz, the capacitanceincreases with glucose concentrations. These frequency-dependent sensorresponses suggest that the glucose-induced permittivity change throughcapacitance can be measured at a fixed excitation frequency.

At the measurement frequencies (lower than 100 kHz), a number ofpolarization mechanisms, such as electronic polarization, ionicpolarization, orientational polarization, and interfacial polarization,can contribute to the measured electrode capacitances. The bindingbetween glucose and PHEAA-ran-PAAPBA can involves complex dielectricchanges. However, conjectural causes of the glucose-dependentcapacitance changes observed in FIG. 30 can occur as follows. First,when glucose permeates through the semi-permeable membrane, it interactswith AAPBA segments in PHEAA-ran-PAAPBA at a two to one ratio to formcyclic ester of boronic acid, resulting in the elimination of twohydroxyl groups. This can cause a decrease of the net permanent dipolemoments in the polymer solution, thereby reducing the ability of thepolymer solution in energy storage and thus the permittivity. Second,glucose binding can lead to variations in the net charge of polymersegments as well as in the polymer conformations, which would alter theelectric double layer structure and result in changes inMaxwell-Wagner-Sillars and counterion polarization. Moreover, thecrosslinking of polymer after glucose binding can increase theviscoelastic resistance of the permanent dipoles on the polymer backboneto alignment with the E-field, leading to a decrease in thepermittivity. As a result, the electrode capacitance exhibited adecrease with glucose concentrations at most of measured frequencies inFIG. 30. Although the underling cause of the crossover of electrodecapacitances at a frequency about 10 kHz requires further investigation,it can be possible due to the glucose-induced changes in the interfacialpolarization effect.

Time-resolved measurements of the differential capacitance in responseto glucose concentration changes were also performed, which allowed forassess the time responses and reversibility of the sensor. For example,the glucose concentration was initially allowed to be equilibrated at 50mg/dL in the test cell and the microchambers. Next, the solution in thetest cell was replaced with another glucose solution at 100 mg/dL. Whenthe glucose concentration inside the microchambers had equilibrated to100 mg/dL, the reverse process was initiated, in which the test cell wasrefilled with a glucose solution at 50 mg/dL. The process of solutionrefilling of the test cell lasted about 10 s, which was sufficientlyfast when compared with the glucose concentration equilibration.

From the 1 data (FIG. 31), it can be seen that, as the glucoseconcentration varies from 50 to 100 mg/dL, the differential sensorcapacitance decreases with time, corresponding to a decrease in thepermittivity of the polymer solution. The capacitance finally saturatesto a constant level, reflecting that the process of glucose permeationand binding has reached a dynamic equilibrium. The time constant of thisprocess was approximately 2.6 minutes. In the reverse process, theglucose concentration in the test cell decreased from 100 to 50 mg/dL.The capacitance increases with time due to an increase of thepermittivity of the sensing polymer solution. The time constant of thereverse process was approximately 3.8 minutes. The longer time constantof the reverse process can be due to the smaller diffusivity of glucosemolecules in the initially more viscous polymer solution. Note thatthese time constants are comparable with the response times ofcommercial systems, which range from 5 to 15 minutes, and can be furtherimproved by shortening the distance between the semi-permeable membraneand the electrodes.

The reversibility of the device responses can be obtained by comparingdifferences in sensor outputs between two separated measurements at thesame glucose concentration. For example, as shown in FIG. 31, thedifferential sensor capacitance varies from 38.633 (averaged over theperiod between 6 and 11 minutes) to 38.626 pF (averaged over the periodbetween 40 and 50 minutes). The difference between the differentialcapacitance over the two periods is about 7 fF or 180 ppm. Thisreversibility was achieved without delicate temperature control as thesingle-module dielectric sensors, but solely depended on differentialmeasurements to compensate for the environmental disturbances,indicating that the differential dielectric sensor can be applied forlong-term, implantable CGM.

The glucose response of the sensor was further assessed by sequentiallyexposing the device to physiologically relevant glucose concentrationsfrom 50 to 500 mg/dL. The measurement started with glucose-free sensingand reference polymer solutions. After the differential capacitancebecame stable, glucose solutions at escalated concentrations werequickly introduced into the test cell of the sensor. From FIG. 32, thesensor capacitance decreases steadily with the glucose concentrationfrom 38.88 to 37.74 pF, indicating a measurement resolution of 0.002mg/dL with an accuracy of 1.75 mg/dL. The differential capacitance wasobserved to have the tendency to become saturated at higher glucoseconcentrations. This indicates that the relationships betweendifferential sensor capacitances and glucose concentrations arenonlinear, and can be represented by a quadratic equation, which isuseful in the in-vivo sensor calibration.

The drift of the sensor output was investigated by exposing the sensorto 50 mg/dL glucose solution over a long period. During this period, theCDC was continuously switched between the differential and thesingle-module measurements. From FIG. 33, the differential capacitanceis steady at 38.62 pf over a period about 4 h. In contrast, asignificant drift about 0.2 pF/h is observed in the single-modulecapacitance, which varies from 57.68 to 56.9 pF. The drift in the singlemodule measurement is possibly due to environmental variations andosmotic pressure, which have been mostly compensated by the differentialmeasurement. These results indicate that the differential measurementeffectively resists the drift in the sensor output and exhibitsexcellent stability that is suitable for long-term CGM.

The sensor stability in face of sudden temperature variations wasobtained in both differential and single-module measurements todemonstrate the ability of the sensor in rejecting temperaturefluctuations, which generally exist in in-vivo environments. To simulatethe actual implantation environment, the temperature of the device wasaltered among physiological temperatures under a closed-loop controlledheating system. As can be seen from FIG. 34, the differential sensorcapacitance has significantly less change and thus, better temperaturestability as compared with the single-module capacitance. As the devicetemperature is changed from 35 to 40° C., the differential capacitancechanges by 0.15 pF, corresponding to a 1.3 pF change in thesingle-module capacitance. This result indicates that differentialmeasurements effectively compensate the influence from temperaturevariations. Here, the compensation of the differential measurement isonly partial, as the temperature-induced capacitance changes are reducedrather than completely eliminated. This can be explained by the mismatchof the thermal-electric properties between the sensing and the referencepolymer and inconsistence during the fabrication of the sensing and thereference electrodes, leading to different capacitive responses upontemperature changes. The temperature stability can be further improvedby careful selection of the reference polymer to achieve a similarthermal-electric property as the sensing polymer.

The device was characterized in-vivo with three sedated laboratory mice.FIG. 35A is a picture showing sensor implantation in one such sedatedlaboratory mouse. The glucose sensors implanted in the subcutaneoustissue of the sedated mice measured the glucose concentrations in ISFcontinuously, while a commercial glucometer sampled the blood sugarlevels in the tail of the mice every 5 minutes.

After sensor implantation, the devices were initiated for 10 to 30minutes to allow the equilibrium of glucose and saline in microchamberswith the environmental ISF. During this process, glucose in the ISFpermeated through the semi-permeable membrane and diffused into thepolymer solutions, which were originally free of glucose beforeimplantation. Simultaneously, the differences in the saline compositionbetween the polymer solutions and ISF were also eliminated through ionexchanges. An exemplary sensor response for this setup process was shownin FIG. 35B, in which the differential sensor capacitance decreases overtime from 0.78 to 0.68 pF, indicating a decrease in the permittivity ofthe PHEAA-ran-PAAPBA polymer solution due to affinity binding betweenthe glucose and the sensing polymer. The differential capacitance iseventually leveled, indicating the completion of the sensorinitialization and the readiness of implantation measurements.

Both the differential sensor outputs and the glucometer readings wererecorded after device initialization. During measurements, blood glucoselevels were first allowed to be purely managed by the metabolism of thesedated mice without intervention from glucose or insulin injections.The glucose levels of mice were then reduced to hypoglycemia throughinsulin injections, and, afterward, increased to hyperglycemia viaglucose injections. The glucometer readings and the changes of thedifferential capacitance calculated with respect to the value at thetime of the first glucometer reading are shown in FIG. 36 for all testedmice. It can be seen that the device output closely followed thecommercial glucometer readings as the blood glucose levels vary over themeasurement periods ranging from 90 to 150 minutes. Time lags betweenthe differential capacitance and the glucometer readings when glucoselevels undergo rapidly changes exist for all tested mice. The lags,which range from 5 to 15 minutes, depend on individual tested subjectsas well as the response times of the tested glucose sensors.

The differential capacitance of the sensor can be calibrated with thereference glucose values obtained from the glucometer using a previouslyintroduced six-point calibration method, as discussed in Example 1. Theclinical accuracy of the a, as compared to G₁ can be quantified using aClarke error grid (FIG. 37), which has five zones labeled with lettersfrom A to E, representing different levels of measurement accuracy. All61 measured points were calibrated and the corresponding a, wasobtained. Here, all points fall exclusively fall into Zone A (83.6%) andZone B (16.4%), whit no point falling into other zones. These resultsindicate that the differential dielectric sensors achieve clinicalaccuracy and good consistency with the glucometer. The test results inthis Example demonstrate that this differential dielectric sensor can beused for subcutaneously implanted devices for long-term, stable, andreliable CGM in diabetes management.

Example 5 An Implantable Monitor Including a MEMS Affinity Sensor and aWireless Interface

In this example, active telemetry is used to construct a wirelessinterface for the implanted glucose sensors. A photograph of a wirelessinterface 3800 and an external reader 3880 are shown in FIG. 38B.Although the dimension of this wireless interface is still relativelarge for fully implantable applications, it allows for verification ofthe feasibility of this wireless interface, and provides a convenientand reliable tool for in-vivo animal studies.

As schematically illustrated in FIG. 38A, the wireless interface 3800contains a low frequency passive transponder 3810 (Texas Instruments,TMS37157), a microcontroller 3830 (Microchip technology, PIC16LF1829), aCDC 3840 (Analog Devices, AD7746), a voltage regulator (TexasInstruments, TPS76901), and a rechargeable battery 3820. Integrated witha glucose affinity sensor 100, this wireless interface can record andsend out a 24 bits sensor capacitance value every 3 seconds. While notreading the data from the transponder, an external reader 3880 (TexasInstruments) provides a continuous RF signal to charge the battery 3820.The CDC 3840 can digitize the capacitance of the glucose sensors, andstore the results into the EEPROM of the CDC. The saved data is thenread via an I2C bus by the microcontroller 3830, which is then sent thedata to the transponder 3810 through a SPI bus. The transponder 3810modulates the digitized data by using two carry frequencies to represent“1” and “0”. In addition, the transponder 3810 can harvest RF power fromthe external reader and provide a voltage at 3.6 V to the rechargeablebattery 3820. All integrated circuit (IC) chips in the wirelessinterface feature low power consumption and are able to enter fromworking mode to sleep mode to save the energy when no operation isconducted.

The functioning of this wireless interface was tested as follows. Asingle-module dielectric affinity glucose sensor, such as described inExample 4, was coupled into the wireless interface 3800. This sensorcontains a parallel plate capacitor formed by a perforated electrode anda bottom electrode. The sensor was original exposed in air, and thecapacitance measured by the wireless interface was approximately 0.76 pF(FIG. 39). Then, the dielectrics sandwiched between the electrodes waschanged from air to water. As a result, the sensor capacitance increasedto 0.9 pF, indicating an increase in permittivity of the sensordielectrics. This result demonstrates that the wireless interface hassuccessfully measured the permittivity-induced capacitance change andcan be applied to implantable glucose detection. In addition, the sensorcapacitance measured by the interface is very stable in water,suggesting similar sensor stability in polymer solutions.

The foregoing merely illustrates the principles of the disclosed subjectmatter. Various modifications and alterations to the describedembodiments will be apparent to those skilled in the art in view of theinventors' teachings herein. Features of existing methods can beintegrated into the methods of the exemplary embodiments of thedisclosed subject matter or a similar method. It will thus beappreciated that those skilled in the art will be able to devisenumerous methods which, although not explicitly shown or describedherein, embody the principles of the disclosed subject matter and arethus within its spirit and scope.

The invention claimed is:
 1. A microdevice for monitoring a targetanalyte in a sample using a polymer capable of binding to the targetanalyte, the microdevice comprising: a semi-permeable membranestructure; a substrate; and a first microchamber and a secondmicrochamber, each formed between the semi-permeable membrane structureand the substrate, each of the first and second microchambers comprisinga suspended element positioned to be spaced apart from the substrate;wherein the suspended element in each of the first and secondmicrochambers comprises a perforated top electrode and the substratecomprises a bottom electrode; wherein the first microchamber is adaptedto receive a solution including the polymer; wherein the secondmicrochamber is adapted to receive a reference solution for screeningeffects not caused by the target analyte; wherein perforations in thetop electrode in the first microchamber are configured to allow thesolution to fill a gap between the top electrode in the firstmicrochamber and the bottom electrode; wherein perforations in the topelectrode in the second microchamber are configured to allow thereference solution to fill a gap between the top electrode in the secondmicrochamber and the bottom electrode; wherein the semi-permeablemembrane structure is permeable to the target analyte and impermeable tothe polymer, thereby when the sample is placed in contact with thesemi-permeable membrane structure, the target analyte, if present in thesample, permeates the semi-permeable membrane structure and enters thefirst microchamber and the second microchamber, respectively, and thepolymer is prevented from escaping from the first microchamber throughthe semi-permeable membrane structure.
 2. The microdevice of claim 1,wherein the binding of the polymer with the target analyte causes achange in the permittivity of the polymer solution.
 3. The microdeviceof claim 1, wherein the top electrode in each of the first microchamberand the second microchamber is supported by at least one anti-stictionpost formed from the substrate.
 4. The microdevice of claim 1, whereinthe semi-permeable membrane structure includes at least twosemi-permeable membrane portions each forming a cover for the firstmicrochamber and the second microchamber, respectively.
 5. Themicrodevice of claim 1, further including the polymer solution in thefirst microchamber and the reference solution in the secondmicrochamber.
 6. The microdevice of claim 5, further comprising acapacitive sensor coupled with each of the first microchamber and thesecond microchamber, the capacitive sensors configured to detect thedifference in capacitance between (A) a first capacitor formed by thesuspended element, the substrate, and the polymer solution sandwichedtherebetween in the first microchamber; and (B) a second capacitorformed by the suspended element, the substrate, and the referencesolution sandwiched therebetween in the second microchamber.
 7. Themicrodevice of claim 1, wherein the perforated electrode in the firstmicrochamber is supported by at least one post formed from thesubstrate.
 8. The microdevice of claim 1, wherein the polymer comprisesa plurality of boronic acid moieties.
 9. The microdevice of claim 8,wherein the polymer comprisespoly(N-hydroxyethylacrylamide-ran-3-acrylamidophenylboronic acid)(PHEA-ran-PAAPBA).
 10. The microdevice of any of claim 1, wherein thepolymer reversibly binds with the target analyte.
 11. The microdevice ofclaim 1, wherein the analyte is glucose.
 12. The microdevice of claim 1,adapted to be implantable in a subcutaneous tissue of a subject.
 13. Themicrodevice of claim 1, further comprising a microheater.
 14. Themicrodevice of claim 1, further comprising a temperature sensor.
 15. Animplantable monitor for monitoring a target analyte in the interstitialfluid of a subject, comprising the microdevice of claim 1 coupled with awireless interface.
 16. The implantable monitor of claim 15, wherein thewireless interface comprises: a capacitance digital converter coupledwith the microdevice and adapted to produce a digital signalrepresenting a measurement of the target analyte in the interstitial ofthe subject; a microcontroller coupled with the capacitance digitalconverter; and a transponder coupled with the microcontroller totransmit the digital signal received from the capacitance digitalconverter to an external reader.
 17. A microdevice for monitoring atarget analyte in a sample using a polymer capable of binding to thetarget analyte, the microdevice comprising: a semi-permeable membranestructure; a substrate; a first microchamber and a second microchamber,each formed between the semi-permeable membrane structure and thesubstrate, each of the first and second microchambers comprising: asuspended element positioned to be spaced apart from the substrate; anda capacitor formed by a top electrode included in the suspended elementand a bottom electrode included in the substrate; wherein the firstmicrochamber is adapted to receive a solution including the polymer;wherein the second microchamber is adapted to receive a referencesolution for screening effects not caused by the target analyte; whereinperforations in the top electrode in the first microchamber areconfigured to allow the solution to fill a gap between the top electrodein the first microchamber and the bottom electrode; wherein perforationsin the top electrode in the second microchamber are configured to allowthe reference solution to fill a gap between the top electrode in thesecond microchamber and the bottom electrode; wherein the semi-permeablemembrane structure is permeable to the target analyte and impermeable tothe polymer, thereby when the sample is placed in contact with thesemi-permeable membrane structure, the target analyte, if present in thesample, permeates the semi-permeable membrane structure and enters thefirst microchamber and the second microchamber, respectively, and thepolymer is prevented from escaping from the first microchamber throughthe semi-permeable membrane structure.
 18. The microdevice of claim 17,wherein the top electrode in each of the first and second microchambersis supported by at least one post formed from the substrate.
 19. Themicrodevice of claim 17, wherein the binding of the polymer with thetarget analyte causes a change in the permittivity of the polymersolution.
 20. The microdevice of claim 17, further comprising thepolymer solution in the first microchamber and the reference solution inthe second microchamber.
 21. The microdevice of claim 17, wherein theanalyte is glucose.
 22. The microdevice of claim 17, wherein themicrodevice is adapted to be implantable in a subcutaneous tissue of asubject.
 23. An implantable monitor for monitoring the target analyte inthe interstitial fluid of a subject, comprising a microdevice of claim17 coupled with a wireless interface.